Since the advent of implantable medical devices, the fields of medicine and electronics have advanced significantly. Accordingly, both the variety and availability of implantable medical devices have increased rapidly. Implantable medical devices to which this invention could be applied currently include not only pacemakers, but also defibrillators, neural stimulators, drug and other therapy delivery systems, monitoring devices some having implanted sensors, and implantable cardioverters, and combinations of these devices, among others.
Early implantable devices were non-invasively controlled to perform simple functions such as turning the device on or off, or adjusting a fixed pacing rate, and even some reporting out of data. These early devices commonly used a magnetic reed switch as a simple telemetry control device. The reed switch could be opened or closed by a magnet held to the patient's skin over the implanted device. In this manner, using the reed switch, the implanted device could be turned on or off. Fixed pacing rates could also be adjusted, for example, by the amount of time the reed switch remained closed, or the number of times it was closed and opened in a given period or a rate or openings and closings, for examples. Such controls, using the reed switch, allowed only for very basic transmittal of signals to implanted devices, and were very limited in function.
As implantable medical devices became more and more sophisticated, the amount of information desired to be transmitted to (and from) the implanted devices grew. Using pacemakers as an example, these functions include switching the device operation into different operating modes, and varying the pacing energy levels delivered to the heart. Pacing rates, device calibration information, and other device parameter settings, and even programming differing therapies and monitoring different body conditions or changing the type of monitoring behavior can also be included in a list of items alterable by and employing the telemetry function. Due to the need to communicate increased amounts of information to implanted medical devices, reed switches became insufficient to adequately perform programming and control functions. As implantable medical devices increased in complexity, it also became desirable for an external programming unit to be able to receive more and better information transmitted by the implanted device. Information that is desirable to be received from an implanted device may include operational status information, information sensed by the implanted medical device such as ECG signals for analysis by a physician, and the like.
There are many telemetry systems that have been used or have been disclosed and these include disclosures on various aspects of such systems particularly suited to implantable devices they are described with reference to in the following U.S. Pat. Nos., which are hereby incorporated by reference:
4,026,305; Brownlee et al., Low Current Telemetry System For Cardiac Pacers PA1 4,267,843; McDonald et al.; Means to Inhibit a Digital Cardiac Pacemaker PA1 4,273,133; Hartlaub et al; Programmable Digital Cardiac Pacemaker with . . . PA1 4,361,153; Slocum et al; Implant Telemetry System PA1 4,373,527; Fischell; Implantable, Programmable Medication Infusion System PA1 4,401.120; Hartlaub et al.; Digital Cardiac pacemaker with Program Acceptance Indicator PA1 4,440,173; Hudziak et al.; Programmable Body Stimulation System PA1 4,515,159; McDonald et al.; Digital Cardiac Pacemaker with Rate Limit Means PA1 4,809,697; Causey, III et al.; Interactive Programming and Diagnostic System for Use With Implantable pacemaker PA1 4,979,506; Silvian; Self-Test System and Method for External Programming Device PA1 5,058,581; Silvian; Telemetry Apparatus and Method for Implantable Tissue Stim . . . PA1 5,107,833; Barsness; Telemetry Gain Adjustment Algorithm and Signal Strength Indication in a Noisy Environment PA1 5,127,404; Wyborny et al., Telemetry Format for Implanted Medical Device PA1 5,137,022; Henery; Synchronous Telemetry System and Method for an Implantable . . . PA1 5,168,871; Grevious; 5,168,871; Method and Apparatus for Processing Quasi Transient Telemetry Signals in Noisy Environments PA1 5,241,961; Henry; Synchronous Telemetry Receiver and Receiving Method for an Implantable Medical Device PA1 5,292,343; Blanchette et al.; Hand Shake For Implanted Medical Device Telemetry PA1 5,324,315 Grevious; Closed Loop Downlink Telemetry and Method for Implantable Medical Device PA1 5,344,431; Merritt et al., Method and Apparatus for Determination of End-Of-Service For Implantable Devices PA1 5,354,319; Wyborny, et al., Telemetry System for an Implantable Medical Device PA1 5,350,411; Ryan et al.; Pacemaker Telemetry System PA1 5,562,714; Grevious; Magnetic Field Strength Regulator for Implant PA1 5,569,307 Schulman et al.; Implantable Cochlear Stimulator having Backtelemetry Handshake Signal PA1 5,693,076; Kaemerer; Compressed Patient Narrative Storage and Full Text Reconstruction From Implantable Medical Devices PA1 A is the coil area in square meters, PA1 F=signal frequency in Hertz, .lambda.=freespace wavelength in meters PA1 Q.sub.s is the coil's in-circuit loaded quality factor PA1 R.sub.S is the total in-circuit loss seen by the coil PA1 I.sub.in is the coil current in amperes, and PA1 R is the link range in meters.
The above-mentioned patents are all incorporated hereinto by this reference. They describe telemetry problems and solutions that run the gamut from dealing with reed switch/device interaction, to protocol for communicating via telemetry, to assisting in the programming of the device, to making adjustments for signal strength. While the list is not exhaustive it is believed to roughly represent the state of the art. Of most relevance to this invention are the ones concerned with adaptation to signal strength, but the others do provide useful background. For this invention we are most concerned with antenna configuration and use, and the prior art in this area is not well developed for the implantable device field.
Typical telemetry systems for implantable medical devices operate with a radio-frequency (RF) transmitter and receiver in both the device and in the external programming unit. The device transmitter and receiver typically use a wire coil element for receiving telemetry signals. The telemetry signals are received from a hand held programming head which is connected to the programming unit, and which is positioned over the implant site in close proximity to the implant. The device transmitter communicates signals to the external programming unit through the programming head. Signals transmitted to the implantable medical device in this fashion are referred to as downlink or downlink telemetry. Signals transmitted from the implantable medical device to the external unit are referred to as uplink or uplink telemetry. The usable uplink range of implantable medical devices is usually quite limited due to the small current amplitudes that the uplink transmitters have to work with.
Implanted medical devices typically have an internal battery(although tapping into body provided energy sources is not unheard of), which provides a power source to drive the normal operative functions of the implanted medical device. When a telemetry session is initiated, this battery must also provide power to operate the RF uplink and downlink telemetry circuitry within the device. This battery has an output terminal voltage, which is limited to a small value of typically 2 or 3 volts. Charge pumping a capacitor can alleviate some of this limitation, but the inventor believes that would be a more costly solution and in any event is not discussed here. During the telemetry system, the battery must also supply a DC current from its terminal voltage to continue the normal implanted device circutry operations.
The power delivered from the battery is calculated as the product of the battery's terminal voltage multiplied by the current delivered from the battery. Since the normal operating functions of the internal medical device circuitry cannot be compromised during a telemetry session, the battery power taken to operate the telemetry circuitry must be limited. The telemetry transmitter and the normal operating functional circuitry both operate at the same battery terminal voltage. Consequently, the only way to restrict the battery power consumed for telemetry is to restrict the amount of current used by the telemetry transmitter and its coil antenna.
The amount of battery current taken for telemetry, while minimized, cannot be set to an arbitrary small value because the telemetry system must also operate over a specified minimum uplink range. In general, from the perspective of this invention, the larger the sweet spot that can be provided with the same amount of battery (or other) current, the better.
The range over which a telemetry link provides acceptable signal transfer, especially for the uplink from the implantable device to the programmer is a key factor in the telemetry system design. Other factors being equal, range depends upon the current amplitude delivered to the coil antenna. Any limitation on the amount of current supplied to the IMD telemetry transmitter and coil antenna will directly limit the telemetry uplink range unless the current is used efficiently within the transmitter and coil antenna. The greater the telemetry transmitter and coil antenna current, the greater the telemetry uplink range becomes, at the expense of less current available for the IMD circuitry.
For our purposes, we prefer to do telemetric communication from our implanted device using the near H field from the coil antenna rather than the E field. This is because the H field wave impedance is much less than it is for E field wave impedance, allowing lower loss signal transmission through the metal can which usually forms the shell of the implanted device, and through the flesh and skin. (The near field is generally considered to be less than 1/6th of the wave length of the carrier wave.) Therefore, in our preferred telemetry, uplink telemetry range depends upon the near field magnetic field strength or amplitude for our preferred embodiments. (For devices ata a range close to the body, or having more uplink transmitter power, or with antenna(e) outside of a metal capsule or shell, the E field may be used as well, and then, the electric field strength would be the limiting factor with respect to range. There are other problems with regulatory authorities when using E fields, another reason we prefer the H field as the data transmission vector).
The magnetic field strength depends on the number of coil turns in the antenna, the area of the coil, and the coil current. The uplink transmitter efficiency depends on the coil quality factor, Q. To increase the telemetry uplink range, the magnetic field intensity must be maintained at an increased distance from the implanted device. The magnetic field may be increased by: adding more turns to the coil; making the coil antenna larger in area, winding it with a larger radius; or by driving the coil with a larger coil current. The larger the coil Q, the more efficient the uplink transmitter circuit becomes
It should be noted that, for either uplink or downlink, it would be desirable to utilize only near-field magnetic fields (H fields), which do not require federal licensing since their amplitude falls off so rapidly with link range. Second, it would be desirable to use these near-fields with the within described inventive coil antenna implementation to increase the uplink range of the telemetry by maximizing the transmitted magnetic field strength. This inventive antenna implementatioin provides this range increase by directly dealing with the problem of the limited fixed voltage supplied by the implantable device's battery.
For a fixed coil voltage, adding more turns to the coil will increase the magnetic field intensity generated by the coil. Unfortunately, adding more turns to the coil will increase the coil's self-inductance and its associated input reactance with the result that less current can then be driven into the coil. The resulting coil drive current is then limited because of the fixed (battery) voltage since the coil antenna's inductance and therefore its inductive reactance increase correspondingly with the number of turns in the coil. The magnitude of current flowing in the coil is equal to the battery voltage divided by the inductive reactance of the coil antenna. Therefore, for a fixed coil voltage, adding more turns to the coil actually results in reduced coil current and smaller magnetic field intensity.
Coil antennas with larger areas may be used to increase the magnetic field intensity. The ideal coil antenna design for uplink applications would have an area as large as can be accommodated, either within, or external to the Implantable Medical Device(IMD) can. Unfortunately, the area of the coil antenna is usually not arbitrary. The coil must be small in area to fit within the volume allotted to it by the overall IMD configuration. The coil antenna should have as large an area as possible in order to produce a large uplink sweet spot. Coil antennas with larger areas produce larger sweet spot volumes. The physical configuration of the antenna essentially determines the sweet spot radius and associated volume.
The sweet spot, at a specified range, for uplink telemetry is defined to be a volume above and adjacent to the site of an implanted medical device. Within the sweet spot volume, at the range specified, the telemetry signal voltage received by the external communicating or programming head equals or exceeds a specified value. The received signal voltage is generated by the time varying magnetic flux density within the sweet spot. (The specified value of signal voltage is defined to be an amplitude greater than the lower limit of signal voltage amplitude the expected to be used for reception of the signal). Larger sweet spot volumes are desired in medical telemetry because placement of the programming head within a large volume is simple to acheive and in an operating room or clinical environment. Thus placement of the programming head relative to the implanted device location will not be critical to maintain a strong telemetry uplink signal, because of the relatively large sweet spot volume acheivable with this invention.
The IMD coil antenna should have a Q as large as possible in order to maximize transmitter efficiency. Too large a Q value will also limit the IMD transmitter's 3-dB bandwidth(BW). For a specified transmitter signal bandwidth, the coils' loaded in-circuit Q is defined as, Q=F/BW, (where F is frequency). The loaded, in-circuit coil Q must provide the required signal bandwidth or some of the transmitted uplink telemetry signal power will be lost. So there are limitations to be placed upon the loaded Q value for the IMD coil antenna. Its in-circuit, loaded Q cannot be made arbitrarily large without directly impacting the transmitter signal bandwidth.
The current made available for telemetry is limited by two factors. First, the limited battery voltage available for a telemetry transmitter and coil antenna in an implantable medical device limits the current available to drive the telemetry circuitry. Second, as explained above, the inductive reactance of the coil antenna will limit coil antenna input current. For a fixed voltage driving the coil antenna, at a fixed frequency, as the coil's inductive reactance increases, the coil's input current is reduced. Also, as the life cycle of the IMD battery progresses, the internal resistance of the battery increases. This increased internal resistance further limits the current available from the battery to drive the antenna coil, since the effect of increased internal resistance is to reduce the available terminal voltage from the battery.
A number of potential solutions to the problems of increasing uplink telemetry range, but which do not involve increasing the coil current have been proposed. These methods however have not proven effective.
Larger inductor coil areas would increase the uplink range when driven by small telemetry currents, but usually are not feasible given the small area allotted to the coil in an implantable medical device. The space allocated for the telemetry circuitry and its associated coil antenna is often limited by the volume remaining after the device circuit components (and the battery) have been placed within the implant can.
An increased number of coil turns would provide a larger magnetic field strength to carry the telemetry signals further. However, since the IMD battery voltage is fixed more coil turns will actually limit the coil current because an increased number of coil turns increases the coil's self inductance, and the inductive reactance of the coil antenna, thereby limiting the coil current.
The loaded coil Q factor should also be maintained at the highest possible level to improve transmitter efficiency, but its value must be consistent with maintaining the specified uplink signal bandwidth.
Given the relatively small volume allotted to the coil antenna, fixed battery terminal voltage, self inductance and inductive reactance limitations of the present telemetry antenna technology, it would be desirable to increase the coil transmit current while maintaining the coil quality factor while remaining within size constraints imposed by implantable medical device cases, to increase the uplink telemetry range. The prior art has not provided an acceptable solution for increasing the telemetry range of an implantable medical device.